Embodiments of the present invention relate generally to diagnostic imaging and, more particularly, to an apparatus and method of acquiring imaging data at more than one energy range using a multi-energy imaging source.
Typically, in computed tomography (CT) imaging systems, an x-ray source emits a fan-shaped or cone-shaped beam toward a subject or object, such as a patient or a piece of luggage. Hereinafter, the terms “subject” and “object” shall include anything capable of being imaged. The beam, after being attenuated by the subject, impinges upon an array of radiation detectors. The intensity of the attenuated beam radiation received at the detector array is typically dependent upon the attenuation of the x-ray beam by the subject. Each detector element of the detector array produces a separate electrical signal indicative of the attenuated beam received by each detector element. The electrical signals are transmitted to a data processing system for analysis, which ultimately produces an image.
Generally, the x-ray source and the detector array are rotated about the gantry within an imaging plane and around the subject. X-ray sources typically include x-ray tubes, which emit the x-ray beam at a focal point. X-ray detectors typically include a collimator for collimating x-ray beams received at the detector, a scintillator for converting x-rays to light energy adjacent the collimator, and photodiodes for receiving the light energy from the adjacent scintillator and producing electrical signals therefrom. Typically, each scintillator of a scintillator array converts x-rays to light energy. Each scintillator discharges light energy to a photodiode adjacent thereto. Each photodiode detects the light energy and generates a corresponding electrical signal. The outputs of the photodiodes are then transmitted to the data processing system for image reconstruction.
A CT imaging system may include an energy sensitive (ES), multi-energy (ME), and/or dual-energy (DE) CT imaging system that may be referred to as an ESCT, MECT, and/or DECT imaging system, in order to acquire data for material decomposition or effective Z or monochromatic image estimation. ESCT/MECT/DECT provides energy discrimination. For example, in the absence of object scatter, the system derives the material attenuation at a different energy based on the signal from two relative regions of photon energy from the spectrum: the low-energy and the high-energy portions of the incident x-ray spectrum. In a given energy region relevant to medical CT, two physical processes dominate the x-ray attenuation: (1) Compton scatter and the (2) photoelectric effect. These two processes are sensitive to the photon energy and hence each of the atomic elements has a unique energy sensitive attenuation signature. Therefore, the detected signals from two energy regions provide sufficient information to resolve the energy dependence of the material being imaged. Furthermore, detected signals from the two energy regions provide sufficient information to determine the materials attenuation coefficients in terms of Compton scatter and photoelectric effect. Alternatively, the material attenuation may be expressed as the relative composition of an object composed of two hypothetical materials, or the density and effective atomic number with the scanned object. As understood in the art, using a mathematical change of basis, energy sensitive attenuation can be expressed in terms of two base materials, densities, effective Z number, or as two monochromatic representations having different keV.
Such systems may use a direct conversion detector material in lieu of a scintillator. The ESCT, MECT, and/or DECT imaging system in an example is configured to be responsive to different x-ray spectra. Energy sensitive detectors may be used such that each x-ray photon reaching the detector is recorded with its photon energy. One technique to acquire projection data for material decomposition includes using energy sensitive detectors, such as a CZT or other direct conversion material having electronically pixelated structures or anodes attached thereto. However, such systems typically include additional cost and complexity of operation in order separate and distinguish energy content of each received x-ray photon.
In an alternative, a conventional scintillator-based third-generation CT system may be used to provide energy sensitive measurements. Such systems may acquire projections sequentially at different peak kilovoltage (kVp) operating levels of the x-ray tube, which changes the peak and spectrum of energy of the incident photons comprising the emitted x-ray beams. A principle objective of scanning with two distinctive energy spectra is to obtain diagnostic CT images that enhance information (contrast separation, material specificity, etc.) within the image by utilizing two scans at different polychromatic energy states.
One technique has been proposed to achieve energy sensitive scanning including acquiring two scans at, for instance, 80 kVp and 140 kVp. The two scans may be obtained (1) back-to-back sequentially in time where the scans require two rotations of the gantry around the subject that may be hundreds of milliseconds to seconds apart, (2) interleaved as a function of the rotation angle requiring one rotation around the subject, or (3) using a two tube/two detector system with the tubes/detectors mounted ˜90 degrees apart, as examples.
High frequency, low capacitance generators have made it possible to switch the kVp potential of the high frequency electromagnetic energy projection source on alternating views and interleave datasets. As a result, data for two energy sensitive scans may be obtained in a temporally interleaved fashion rather than with separate scans made several seconds apart or with a two tube/two detector system. In order to improve contrast and reduce or eliminate beam hardening artifacts, it is desirable to increase energy separation between high and low kVp scans. Energy separation may be increased by increasing energy in high kVp scans. However, high kVp scans may be limited due to system stability at high voltage.
Alternatively, energy separation may be increased by decreasing energy in low kVp scans. However, x-ray attenuation may occur for low kVp projections to the extent that system noise may swamp a received signal, and x-ray attenuation typically increases as the size of the imaging object increases. As may be experienced in conventional single kVp imaging, imaging of some objects at, for instance, up to 120 kVp can cause projection data to be contaminated as detected signals become so weak that they are swamped out by other interfering signals such as electronic system noise and scattered x-ray noise. Thus, in conventional CT it is possible to intervene with a low signal mitigation algorithm to avoid low signal streaking artifacts in images. Such algorithms may be applied to one or both sets of scan data in a dual energy application, as well.
However, as understood in the art, low signal mitigation algorithms are typically data smoothing filters that operate along a detector channel, detector row, and/or view dimensions. And, although known algorithms may reduce streaking, they also may reduce high spatial frequency content of data samples, and therefore resolution, in resulting images. Thus, there is a need for low signal mitigation in potentially a large percentage of dual or multi-energy exams that are conducted.
Therefore, it would be desirable to design a mitigation scheme for low kVp imaging that does not compromise high spatial frequency content thereof.